Measurement of volumetric fluid flow and its velocity profile

ABSTRACT

A method to measure spatial fluid flow components and their velocity profiles in a number of locations in a cross-sectional area of a lumen or other body cavity by using ultrasound in which the cross-sectional area is interrogated by a plurality of ultrasonic beams; the estimation of the spatial flow components is obtained from a combination of estimations of axial, lateral and total flow; the estimation of one or more flow components is obtained through any combination of time-shift and decorrelation analysis of two or more beam-signals of the interrogating ultrasound transducer; and the estimation accuracy is further improved by the use of a reference decorrelation curve obtained from experiments or beam-theory or both.

This application is a continuation of Ser. No. 09/297,538 filed Jul. 2,1999, U.S. Pat. No. 6,213,950 which is a 371 of PCT/NL97/00504 filedSep. 2, 1997.

The assessment of fluid flow in the human body is important for medicaldiagnosis. For example, blood velocity and the volumetric flow (i.e.,the volume of blood flowing through a blood vessel, e.g., in liter persecond) routinely assist clinical decisions.

Various ultrasound techniques can be used to measure the motion ofscattering materials such as blood, body fluids and tissue. Ultrasoundcontrast agents can also be used to enhance signals from fluids withinsufficient scatter properties. For example, blood velocity can bemeasured in a small volume using the Doppler principle. In echographicB-scanning, multiple estimates of blood velocity in the plane of thescan can be combined with the gray-scale echo image by colouring.

Miniaturised ultrasound transducers can be placed inside the lumen of avessel or other body cavities to obtain a cross-sectional echo image.The same ultrasound echo signals can be used to measure the velocity ofthe flowing blood or other fluids.

The purpose of the invention is to provide a method for measuringvolumetric fluid flow and its velocity profile in a lumen or other bodycavity. According to the invention an ultrasonic method is provided tomeasure volumetric flow through a lumen by accomplishing simultaneouslyand in situ (in place) the steps of

a) measuring the local velocity of the scattering medium perpendicularto the ultrasound scan plane and

b) integrating such velocity measurements over the area of the lumen.

An approximation to the above method in accordance with the inventioncomprises the steps of

a) calculating the average value of the velocity of the scatteringmedium perpendicular to the ultrasound scan plane,

b) calculation of the area of flow, and

c) multiplying the average velocity by the area to obtain volume flow.

Furthermore, in order to reduce the number of calculations, the averagevelocity can be approximated by measurement of the velocity in asub-region of the area of flow smaller than the total area of flow, andvolume flow can be computed as before.

In the method according to the invention the scattering fluid ofinterest may be blood. For purposes of explaining the invention, theinvention shall be discussed in relation to blood. Other fluids may bemeasured in a similar manner. Blood is composed by red blood cells(RBCs), white blood cells and platelets suspended in a liquid calledplasma. Because the size and density of RBCs is large compared to thatof white cells and platelets, backscatter of blood is attributed to thered blood cells. The measurement of blood velocity comprises the stepsof:

a) obtain (transmit pulse and receive echo) two or more subsequent echosignals from a single (or a slightly changed) position of the ultrasoundtransducer at controlled interval(s) of time Δt,

b) measure one or more displacements of the blood relative to the beam,Δd, and

c) compute velocity from the ratios of displacements over the timeinterval, v=Δd/Δt.

One embodiment of the invention relates to the measurement of volumetricflow and velocity imaging from within the lumen of a blood vessels usingintravascular ultrasound. It has to be mentioned here that conventionalultrasound has already been proposed to measure blood velocity on theplane of the scan. The scan is usually oriented along the blood vessel.If the scan plane is oriented perpendicular to the blood vessel,volumetric flow could also be computed as describe hereinafter.

The invention will be described in details with reference to theaccompanying drawings, in which

FIG. 1 is a schematic view of a lumen in which an ultrasound catheter ispositioned;

FIGS. 2a-2 c are graphs of echo-signals received by an ultrasoundtransducer and of the time shift between echo-pairs;

FIG. 3 is an illustration of the characteristic decorrelation for asingle range;

FIG. 4 is an illustration of the decorrelation “calibrated” velocityestimation procedure;

FIG. 5 is a schematic view of a lumen and the scan plane in such lumen;

FIG. 6 illustrates the direction of flow relative to thethree-dimensional orthogonal axes centered on the transducer aperture;

FIG. 7 is a graph showing velocity profiles at different flowvelocities;

FIG. 8 shows different flow velocity images computed within one heartcycle; and

FIG. 9 is a graph showing, simultaneous calculation of phasiccross-sectional area and phasic volumetric flow.

Local estimates of blood velocity are obtained by means of echo-signaldecorrelation and time-shift analysis. By means of an ultrasoundtransducer, a sound pulse is transmitted into the scattering medium;backscattered echoes from the medium are received by the same (or aseparate transducer) and converted to an electric signal suitable forstorage and processing. In FIG. 1 a rotational scan of the beam of atransducer is depicted.

The velocity of a moving object can be calculated by measuring thedisplacement of the object during a given interval of time. The ratio ofdisplacement and time interval is the velocity.

The displacement of an ultrasound-scattering material (such as blood)moving through the beam of an ultrasound transducer results inconcomitant changes in the received echo signal. For example, FIG. 2ashows a sequence of five echo signals (S1 through S5) obtained in anexperiment where a scattering material is progressively displacedthrough an ultrasound beam. It can be observed that with increasingdisplacement, the echo signals progressively a) shift (advance) in time;and b) change in shape (FIG. 1a).

The correlation coefficient ρ is a measure of the similarity (ordissimilarity) between a pair of signals and is defined such that ρ=1(100%) when there is total similarity, and ρ=0 when the signals are noterelated at all. A decrease in correlation is termed decorrelation.

In an experimental example, the four correlation coefficients (or simplycorrelation for short) computed between echo-signal pairs S1-S2, S1-S3,S1-S4, and S1-S5 are shown in FIG. 2b. Correspondingly, the progressivetime shift between the same echo signal pairs is computed and shown inFIG. 2c, In this experiment the time interval was Δt=250 μs. In generalthe time interval must be sufficiently short to warrant recognition ofthe advance in time and shape change of the echo signals: that is, ifthe time intervals is long relative to the velocity, the echo signalwill change drastically precluding measurement of time shift anddecorrelation.

Echo decorrelation is mainly a function of the beam characteristics (thewidth of the beam among others). For example, for a beam with beam widthof 1 mm, the echo signal would be totally decorrelated after a 1 mmdisplacement of the scattering medium; however, a different transducerwith a beam width of 2 mm would maintain some of the correlation after a1 mm displacement since the scatterers are still within the beam width.

Beam characteristics are range dependent. Consequently, an ultrasoundbeam exhibits a range-dependent decorrelation characteristic. Byexperimentally or theoretically assessing the decorrelation for atransducer at all ranges and for displacements in all directions weobtain what we can call the “characteristic decorrelation” of the beam.Once the characteristic decorrelation has been assessed, measureddecorrelations in blood or tissue can be converted to displacement. Forexample, for a given range and direction of displacement across aparticular beam, a characteristic decorrelation curve is illustrated inFIG. 3; in this example, a measured decorrelation value of ρ₁ wouldcorrespond to a displacement of d=0.1 mm.

Thus, the decorrelation characteristic of the transducer serves thepurpose of a calibration factor which can be used to convert measureddecorrelations into displacement and velocity.

One or more echo decorrelation values can be involved in the computationof a velocity value. Using a single decorrelation value between a singlepair of echo signals, the velocity is computed as the ratio of thedisplacement obtained from the characteristic decorrelation at a giventime interval. Using the example in FIG. 3, of the time interval betweenthe echo acquisitions giving rise to the decorrelation of value ρ₁ wasΔT=1 ms, then the velocity v would be d/ΔT=0.01 mm/0.001 s=10 cm/s. Inpractice the decorrelation characteristic versus displacement functionmay not be easily described analytically i.e., by a formula or theformula may not be suitable for inversion (obtain displacement fromdecorrelation). However, pre-calculated values of characteristicdecorrelation for small increments in displacement can be arranged in a“look-up” table: then, by “looking up” the decorrelation value in thetable, the corresponding displacement can be obtained. This is termedthe look-up-table (LUT) method. The LUT method makes no assumptionregarding the shape of the characteristic decorrelation.

Estimation of velocity based on a single decorrelation value is verysensitive to errors caused by other sources of decorrelation that arenot related to motion. However, by using the rate of change from two ormore echo decorrelation values, improved accuracy and precision ofestimates of velocity may be possible. Although the improvement inprecision can be expected since there is an implicit averaging proceduretaking place, the improved accuracy is specific to this application.This particular scheme for averaging multiple decorrelation estimates ismentioned as an example and is not intended to exclude the possibilityto use other averaging procedures with similar improvement.

Curve-fitting algorithms can be applied where a curve is fitted over anumber of decorrelation measurement: the simplest curve being a straightline.

The linear fit method is illustrated in FIG. 4, where

a) a straight line is fitted over two or more decorrelation values.Since, by definition the decorrelation must be 1 for zero displacement,a single parameter (i.e., the slope of the line, termed thedecorrelation slope), defines the curve formed by decorrelationmeasurements;

b) similarly, another straight line fit is performed in thecorresponding area of the characteristic decorrelation curve; thisyields a characteristic decorrelation rate; and

c) the ratio of the decorrelation slope and the characteristicdecorrelation rate yields the velocity of blood (in units of mm/s).

FIG. 4 shows:

(a) Beam decorrelation estimated by experiment or theory; here shownonly for one distance from the transducer.

(b) Decorrelation measured with ultrasound at five intervals of time(dots) with straight line fit.

(c) Velocity estimate using the decorrelation rate for the transducerfrom (a) and the decorrelation slope from the measurement (b) atcorresponding depths.

When the curve formed by subsequent decorrelation measurement is notwell approximated by a straight line, the linear-fit approach can leadto biased estimation. However, it follows from the above descriptionthat when a linear fit is inappropriate, a higher-order fit can beapplied and more than one parameter is required to describe the bestfit.

The LUT method with multiple decorrelation estimates involves the stepsof

a) obtain from the characteristic decorrelation LUT the displacement foreach measured decorrelation value obtained at subsequent time intervals;and

b) a straight line fit is performed on the displacement versus timeinterval plot. The slope of the straight line fit is the velocity ofblood.

An alternative approach to that of fitting of a particular trend to thedecorrelation measurements is to simply calculate the average of allavailable decorrelation estimates. It is important to recollect that thebasic unit of the correlation algorithm is the cross-product of two echosignals. For a pair of discrete echo signals s₁(i) and s₂(i), thecorrelation coefficient is given by$\rho = {\frac{\sum\limits_{i}{s_{1} \times s_{2}}}{\sqrt{\sum\limits_{i}{s_{1}^{2}{\sum\limits_{i}s_{2}^{2}}}}}.}$

Multiple decorrelation estimates can be made simultaneously by squaringthe sum of a number of echo signals, since such operation yields a sumof cross-products. In order to illustrate this, we establish therelationship between the correlation coefficient and the normalized sumof squares as follows:${S = \frac{\sum\limits_{i}( {s_{1} + s_{2}} )^{2}}{{\sum\limits_{i}s_{1}^{1}} + {\sum\limits_{i}s_{2}^{1}}}},\quad {and}$$S = {\frac{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}} + {2{\sum\limits_{i}{s_{1}s_{2}}}}}{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}} = {1 + {\frac{2{\sum\limits_{i}{s_{1}s_{2}}}}{{\sum\limits_{i}s_{1}^{1}} + {\sum\limits_{i}s_{2}^{1}}}.}}}$

When the two terms in the denominator of the above equation are similar,their arithmetic and geometric mean values can also be assumed to besimilar, thus${1/{2\lbrack {{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}} \rbrack}} \approx {\sqrt{\sum\limits_{i}{s_{1}^{2}{\sum\limits_{i}s_{2}^{2}}}}.}$

Then we can write

S≈1+ρ.

With the above formula, decorrelation can be estimated from thenormalized sum of squares. When the squared sum involves more than twoterms, additional factors weighting the cross-products arise and dependon the number of terms in the square sum. For example, for the squaresum of three echo signals obtained at intervals of time Δt, we get${\sum\limits_{i}( {s_{1} + s_{2} + s_{3}} )^{2}} = {{\sum\limits_{i}\lbrack {s_{1}^{2} + s_{2}^{2} + s_{3}^{2} + {2s_{1}s_{2}} + {2s_{2}s_{3}} + {2s_{1}s_{3}}} \rbrack} \approx {\sum\limits_{i}\lbrack {s_{1}^{2} + s_{2}^{2} + s_{3}^{2} + {4s_{1}s_{2}} + {2s_{1}s_{3}}} \rbrack}}$

since we can assume that${\sum\limits_{i}{s_{1}s_{2}}} \approx {\sum\limits_{i}{s_{2}{s_{3}.}}}$

Thus the correlation for echo signals spaced by a single time intervalhas a weighting factor of four, while for a spacing of two timeintervals the weighting factor is two. This differential weighting ofcorrelation at different time intervals must be taken into account toobtain an accurate estimate of the average decorrelation.

Similarly, decorrelation estimates can be obtained from the squareddifference between echo signal pairs:${D = \frac{\sum\limits_{i}( {s_{1} - s_{2}} )^{2}}{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}}},$

and following similar steps as above, we obtain

D≈1−ρ

With the above formula, decorrelation can be estimated from thenormalized squared difference of echo signals. In order to furthersimplify the calculation requirements, it is possible to substitute theabsolute difference for the squared difference. In this way, theoperation of calculating square values is avoided.

Note that while prior art teaches that decorrelation can be used toassess displacement, the improvement in accuracy due to the combinationof multiple decorrelation values is novel.

For the purpose of measuring volumetric flow, the velocity component ofthe flow normal (perpendicular) to the scan plane must be assessed.Alternatively, the angle between the scan plane and the flow must beassessed.

This is illustrated for the vascular application in FIG. 5. The bloodvelocity must be computed for blood velocity imaging, but the normalflow component must be computed for volumetric flow estimation.

In general, the direction of flow can have any arbitrary angle withrespect to the reference axes of the ultrasound beam. This isillustrated in FIG. 6 where the direction of flow is shown relative tothe three-dimensional orthogonal axes centered on the transduceraperture. These axes in combination with the direction of the scan(sweeping of the beam) give rise to three spatial directions that aretermed axial, lateral, and elevational (FIG. 6). When a scatteringmedium displaces exclusively along the axis or the transducer (axialdisplacement), the motion can be assessed from the time shift of theecho signal when axial decorrelation is sufficiently low. When themotion occurs exclusively across the axial direction of the transducer,the echo signal decorrelates as the original scattering blood particlesmove out of the beam and new blood particles move into the beam.Displacement across the beam and in the scan plane is termed lateraldisplacement. Displacement across the beam and across the scan plane istermed elevational displacement.

Axial, lateral and elevational displacements introduce echodecorrelation giving rise to axial, lateral and elevationaldecorrelation components. Additionally, axial displacement introduces atime shift. Thus, in order to fully characterize displacement withrespect to a scan plane, all these decorrelation components must becomputed. Axial and elevational displacement can be measured asdescribed above. Lateral displacement can be estimated by“cross-decorrelation”, that is decorrelation analysis between-echosignals from adjacent beam locations within the scan plane. Therelationship between these decorrelations must also be assessedexperimentally or theoretically.

In a hypothetical example, assuming a linear decorrelation withdisplacement, lateral and elevational decorrelations can be combined inthe squared sense (that is, the square of the total decorrelation is thesum of the square of the lateral and elevational decorrelations); then,knowing the lateral displacement from cross-decorrelation analysis theelevational component of displacement can be isolated.

Analogously, the presence of axial, lateral and elevational velocitycomponents can be considered. It is important to note that for thepurpose of measuring volumetric flow, the velocity component normal tothe scan plane must be assessed. When the direction of flow is notperpendicular to the scan plane (elevation direction), flow velocityestimated with decorrelation may lead to biased estimation of volumeflow (unless the angle between the direction of flow and the scan planeis assessed).

In the common practice of intravascular ultrasound, the main componentof displacement is in the elevational direction. Thus, the contributionof axial and lateral displacements may be neglected in somecircumstances without significant deterioration of the velocityestimates However, the contribution of lateral and axial componentsshould be kept under consideration. These components can increase duringthe examination of curved vessels and when secondary flow is present,among other possibilities.

Echo signal decorrelations also occurs due to sources that are unrelatedto motion. For example, electronic noise present in the echo signalsresults in decorrelation. Two echo signals can differ only due tocorruption of independent realisations of the noise source. Thus, forimproved estimation of velocity from decorrelation, decorrelation fromsources of non-motion-related decorrelation must be isolated.

By computing the decorrelation of echoes from stationary tissue (e.g.,vessel wall), the decorrelation due to sources other that motion can beassessed. Then, this decorrelation can be deducted from the totaldecorrelation measured from moving blood (under the assumption thatblood and tissue echo signals contain the same amounts of electronic andquantization noise).

The main goal of the invention is to determine volumetric flow andvelocity of blood with respect to the vessel through which it flows.However, the above scheme estimates velocity with respect to the beamunder the assumption that the transducer is fixed in position within thevessel. Normally, the transducer can and will move with the pulsation offlowing blood, adding an undesired motion component. In order to improvevelocity estimation, the motion of the transducer with respect to thevessel wall must also be assessed. Transducer motion can be assessed bymeasuring the time shift of echoes from quasi-static (relative to thehigh velocity blood) vessel wall tissue.

In general, the local velocity of blood varies within the vessel lumen;that is, blood tends to move slower near static features (e.g., thevessel wall, the ultrasound catheter) and faster away from thosefeatures (e.g., towards the center of the free lumen). Thus, to fullycharacterize flow it is important to estimate blood velocity in alocalized manner (note: commonly used Doppler-based techniques for flowestimation are based on the estimation of the peak value of the velocitywithin the lumen and the assumption of a velocity distribution).

Blood motion measurement scan be achieved in a localized manner byanalysis of gated (i.e., for “small” segments of) echo signals whichcorresponds to the local motion of small parts of the blood volume Forexample, FIG. 2 shows gated signals representing echoes from 0.25 mm ofblood. Many such local estimates of blood are measured in adjacentregions along the beam and can be combined to form a “velocity profile”.Velocity profiles obtained at different flow velocities are shown inFIG. 7: note the higher velocities near the center and low velocitiestowards the static wall. Adjacent beam positions in a scan plane can beused to form a map or image velocity. The sweeping of the beam(scanning) can be achieved.

a) by physically moving the transducer,

b) by physically moving a mirror which reflects the beam of a statictransducer, or

c) by electronically generating of the beam using an array oftransducers elements.

Velocity images can be superimposed on the gray-scale image of statictissues by coloring image pixels according to the magnitude of flow. Forexample, from an experiment conducted in a live pig, four flow velocityimages computed within one heart cycle are shown in FIG. 8a. Thus,unlike any previously described method, an image of the velocity ofblood flowing normal to the cross-section of the scan plane is obtained(illustration in FIG. 1, measurement in FIG. 8).

Additionally, velocity components or the volume flow can be convertedinto an audible signal in order to provide a different way to presentflow information. For example, a converter could be used to transformflow data into an analog signal which, following amplification, coulddrive a loudspeaker. This may be a useful feature when the operator isunable to look at the monitor while manipulating the ultrasoundtransducer. For instance, the range and mean value flow velocity couldbe represented by the bandwidth and the pitch of the output sound. Thispresentation of flow information is analogous to what is used in Dopplersystems where the frequency of the Doppler signal is by default in theaudible range and needs only be amplified and connected to a speaker.However, here the sound is synthesized from independently computed flowinformation.

Additionally, decorrelation based estimation of motion can be used todifferentiate areas where there is moving blood (that is, the free lumenarea) from areas where tissue is static. Blood flows at a much highervelocity relative to the motion of healthy or diseased tissue (such asthe vessel wall, plaques and dissections).

The prior art in the intravascular ultrasound application describes thecombined use of two catheters to measure volumetric flow. A firstintravascular ultrasound catheter is used to measure free lumen area andis not capable of measuring flow. A second Doppler-ultrasound catheteris used to assess the velocity of blood flow and is not capable ofmeasuring the area of flow. Additionally, the Doppler catheter cannotmeasure local velocity at many spatial locations: a velocity profilewithin the lumen is not measured but assumed based on a singlemeasurement of velocity (peak or mean).

The combination of measured area and assumed velocity distributionyields volumetric flow. However, the cross-sectional area measurementand the flow estimation are performed at two spatially separatedlocations. Alternatively, cross-sectional area can be measured in onelocation of the vessel, and later the velocity can be measured in thatsame location. Thus, the prior art is limited to either “simultaneous”or “in place” measurements, but does not teach both.

In the current invention integration of the measured velocity map overall points in a cross-section of the free lumen area yields the volumeof blood flowing through the vessel. Since the area of integration andthe flow velocity is computed from the same signals, it is a fact thatthe estimation of area and velocity are simultaneous and in place.

Alternatively, a reduced, but representative number of points within across-section of the free lumen area can be used to estimate the averagevalue of flow velocity in the entire free lumen area. The averagevelocity can be estimated from a partial area of the flow area when thevelocity distribution is known to first approximation. Simultaneously,the area of flow can be calculated at the same cross-section. The areaof flow times the average flow yields the volume of blood flowingthrough the vessel. Like in the Doppler approach, a velocity profilewithin the lumen must be assumed to calculate the average velocity basedon a restricted number of measurements of velocity within the lumen.However, unlike the Doppler approach, in this alternative implementationvelocity and flow are measured simultaneously and in place.

Since this measurement can be performed at regular intervals that aresmall compared to the period of the heart or respiratory cycle, phasicvolumetric flow can be assessed (phasic meaning the time history withina cycle). Phasic volumetric flow measured in a live-pig experiment isshown in FIG. 8b.

The relationship between local blood pressure and the volumetric flowyields additional hemodynamic information of extreme clinical relevance.Particularly, the change of the pressure-flow relationship in responseto vaso-active drugs is used to investigate the reaction of differentparts of the vasculature. For this reason, several methods have beendeveloped that combine pressure-sensing catheters with intravascularDoppler-ultrasound assessment of volumetric flow. Alternatively, lumenarea or diameter can be used as an estimator of pressure and velocitycan be used as a measure of volumetric flow. A limitation common tocurrent multi-variable methods is the inability to assess thehemodynamic variables coincidentally in time and space. Additionally,multiple catheter approaches suffer from possible interference betweendevices: for example, in the combination of intravascular ultrasound forarea measurement and Doppler catheter for velocity measurement, theintravascular catheter disturbs the blood flow and therefore affects thevelocity that is measured by the Doppler catheter.

Since the endoluminal pressure is intimately related to the change incross-sectional lumen area (particularly, there is a linear relationshipbetween pressure and lumen diameter toward late diastole), thisinvention can provide concurrent and coplanar measurements ofcross-sectional lumen area and volumetric flow. This is shown in FIG. 9where phasic cross-sectional area and phasic volumetric flow werecalculated simultaneously and from the same scan plane from the live pigexperiment. The phasic relationship between area and flow can provideinformation on the resistance of the vasculature. Spectral analysis ofphasic area and flow can be used to assess arterial impedance.

So far the measurement of flow has been described assuming that thebackscatter from blood is uniform within the lumen and in time. However,normally red blood cells tend to form clusters, a process calledaggregation, and arrange themselves in “strings” called rouleaux. In thefollowing blood, rouleaux are positioned along the direction of flow.The presence of RBC aggregation and rouleaux is a function of the cyclicvariation of the local shear: regions of high shear (near staticstructures) have low aggregation and areas of low shear (near the centerof the free lumen) have high aggregation. Thus, a spatial distributionof RBC aggregation as well as a cyclic temporal variation with the heartrate are known to exist. In intravascular ultrasound imaging, theseeffects are manifested as a cyclic and spatial variation of the echointensity (since larger aggregates of RBCs also backscatter strongerechoes). Thus, examination of the echo intensity and the backscattercoefficient function can yield information about the scattereraggregation.

The shape and size of the back scattering particles can also have asignificant effect on the decorrelation phenomenon. Clearly, the echofrom a single point scatterer moving across a beam will decorrelatefaster than the echo arising from a long string of aligned scatterersmoving across the same beam. Thus, the dependence of decorrelation onblood backscatter should be compensated for improved accuracy in volumeflow and velocity estimation. For example, long rouleaux or clusterspresent in the central (low shear) part of the free lumen would resultin local underestimation of velocity.

Analogously to the determination of the characteristic decorrelation ofan ultrasound beam, for each shape and size of scatterers we can obtaina “scatterer characteristic decorrelation” function. Then, from thescatterer type/shape can be estimated from backscatter analysis and thevelocity estimates can be compensated for scatterer characteristicdecorrelation.

The effect of RBC aggregation on decorrelation is a function of the sizeof the aggregate relative to the wavelength. Therefore, applicationsthat utilize high frequency ultrasound (i.e., wavelengths similar to theaggregate size) may be expected to experience a higher dependence onaggregation. Conversely, low ultrasonic frequencies (i.e., wavelengthsmuch larger than the RBC diameter) may be expected to be less affectedby aggregation-dependent decorrelation. In practice, reasonableestimates of flow can be obtained without backscatter-dependentcompensation.

What is claimed is:
 1. A method to measure spatial fluid flow componentsand their velocity profiles in a number of locations in across-sectional area of a lumen or other body cavity by using ultrasoundcomprising the steps of: a) interrogating the cross-sectional area witha plurality of ultrasonic beams; and b) estimating the spatial flowcomponents including at least an elevational flow component throughtime-shift and decorrelation analysis of two or more beam signals of theplurality of ultrasonic beams by using a reference decorrelation curve.2. The method of claim 1 herein the reference decorrelation curve isobtained from one of either; experiments or theory.
 3. The method ofclaim 1, further comprising the step of simultaneous and in placeestimation of a flow area.
 4. The method of claim 3, comprising the stepof combining a component of flow velocity normal to a cross-section ofan area of flow in a number of spatial locations in the flow area toderive a momentary value of volume flow.
 5. The method of claim 3,comprising the steps of combining an average over a number of spatiallocations of the component of flow velocity normal to a cross-section ofan area of flow and the flow area to derive a momentary value of volumeflow.
 6. The method of claim 5, wherein a calculation of the averagecomponent of flow velocity normal to a cross-section of a total area offlow is from a partial section of a total area of flow.
 7. The method ofclaim 3, further comprising the steps of combining the component of flowvelocity normal to a cross-section in a number of discrete spatiallocations and the corresponding areas of the discrete spatial locationsto derive a momentary value of volume flow.
 8. The method of claim 7,wherein the step of combining the normal velocity and correspondingdiscrete areas is an integration operation.
 9. The method of claim 1further comprising the step of compensating for an axial decorrelationcomponent of a scattering medium.
 10. The method of claim 1 furthercomprising the step of compensating for non-motion-related decorrelationcomponents.
 11. The method of claim 1 further comprising the step ofcompensating for the velocity measurements for relative motion of atransducer with respect to relatively static tissues.
 12. The method ofclaim 1 wherein the decorrelation analysis further includes squaringdifferences of echo signal calculations.
 13. The method of claim 1wherein an average decorrelation is derived from squared sums of echosignals.
 14. The method of claim 1 wherein a velocity measurement isconverted into an audible sound.
 15. The method claim 1 wherein aninterrogating transducer is an intraluminal ultrasound transducer wherethe beam is scanned in a plane by one of the following: (a) rotating atransducer; (b) rotating a mirror which redirects the beam of a statictransducer; (c) electronically generating the beam through amulti-element transducer without transducer motion; and (d) acombination of (a) and (c).
 16. The method of claim 1 wherein theplurality of ultrasound beams are scanned through adjacent spatiallocations in a sequential step-wise way.
 17. A method for measuring avelocity of blood in a lumen or other body cavity using ultrasoundthrough echo signal decorrelation and time-shift analysis comprising thesteps of: interrogating a cross sectional area of the lumen or otherbody cavity with a plurality of ultrasonic beams; measuringdecorrelation of the blood from two or more echo signals received fromthe ultrasonic beams over a known time interval; converting the measureddecorrelation to a displacement by comparing the measured decorrelationto a characteristic decorrelation; and calculating the velocity of theblood from the displacement value and a known time interval.
 18. Themethod of claim 17 wherein the characteristic decorrelation is adecorrelation curve.
 19. The method of claim 17 wherein thecharacteristic decorrelation is a set of discrete values.